Blood pump

ABSTRACT

A blood pump has a hollow body in which an impeller with a spiral blading produces an axial propulsion of blood along the impeller, as well as an at least partly actively stabilized magnetic bearing device and a hydrodynamic bearing device for the impeller. The impeller may be set into a rotation about a rotation axis of the impeller with a motor stator located outside the hollow body. The hollow body has an inlet for the flow of blood into the hollow body in an inflow direction which is essentially parallel to the rotation axis, and an outlet for the outflow of the blood out of the hollow body in an outflow direction which is offset to the rotation axis of the impeller to produce a non-zero outflow angle (α) between the inflow direction and the outflow direction. A total artificial heart can be formed from two such blood pumps.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. Non-Provisional applicationSer. No. 15/192,285, filed Jun. 24, 2016, the entire contents of whichare incorporated by reference, which is a continuation of U.S.Non-Provisional application Ser. No. 13/505,368, filed Aug. 21, 2012,the entire contents of which are incorporated by reference, which is a371 nationalization of PCT/EP2010/006863, which in turn claims benefitof U.S. Provisional Application 61/258,932 filed Nov. 6, 2009, and ofEuropean application 09075495.3 filed Nov. 6, 2009.

BACKGROUND

The invention relates to the field of blood pumps.

A blood pump here and hereinafter is to be understood as a pump whichserves for supporting or creating a blood flow within a human or animalbody and is suitable for implantation into the thorax of a human oranimal, outside the heart. With left ventricular assist devices (LVAD),a connection exists between the left heart half and an inlet of theblood pump as well as between the outlet of the blood pump and the aortadeparting from the heart, for the support or creation of the bloodcirculation through the body (systemic circulation). With rightventricular assist devices (RVAD) there exists a connection between theright heart half and the pulmonary artery stem which leads to the leftand to the right pulmonary artery, (or a direct connection between theRVAD and the left and/or right pulmonary artery) for the support orcreation of the blood circulation through the lung (pulmonarycirculation). The blood, within the blood pump, is led through a hollowbody which is part of a pump housing or is arranged in such a pumphousing. A rotating impeller, with a blading for producing a pressureand a flow of the blood resulting therefrom, is provided in the hollowbody. So-called total artificial hearts (total heart pumps) contain aleft ventricular and a right ventricular assist device (blood pump) forthe support or the creation of the complete blood circulation. Flexibleconnection tubes or connection pipes as well as, as the case may be,flow bends or elbows, are applied for creating the mentioned connectionsbetween the blood pump and the heart or blood vessels. Moreover, atleast one cable line is necessary for the energy supply and, as the casemay be, for the control of the blood pump, said cable line connectingthe blood pump to an energy storer and, as the case may be, to a controlunit.

A main problem with the implantation and the use of such blood pumps, inparticular total artificial hearts, is the spatial requirement of suchblood pumps and the flexible connection tubings as well as the cableline in the thorax space in the vicinity of the heart.

A further difficulty lies in the danger of destruction of the bloodcells (hemolysis) due to the blood pump, in particular at the mechanicalbearings of the impeller, narrowings and abrupt changes in direction ofthe blood flow through the blood pump, as well as by way of largepressure gradients within the blood pump. With the design of bloodpumps, for this purpose, often mechanical bearings of the impeller arereplaced by a magnetic and/or hydrodynamic bearing.

An additional problem lies in the fact that a significantly smallerblood pressure needs to be created for the pulmonary circulation, thanfor the systemic circulation, wherein however the same blood volume pertime needs to be transported through both blood circulations. The bloodpressure produced by the blood pump depends on the rotational speed ofthe impeller of the blood pump. It has been found to be difficult todesign a pump which is suitable for setting very different values of theblood pressure within a range of about 5 mmHg up to about 150 mmHg witha stable, constant volume flow between 0 l/min to 20 l/m in adapted tothe physiological conditions, and which in this manner may be applied asan RVAD as well as an LVAD, or which is suitable for the design of atotal artificial heart.

It is therefore the object of the present invention, to suggest a bloodpump as well as a total artificial heart, which solves or at leastreduces the problems mentioned above. A corresponding blood pump ortotal artificial heart should thus have an as low as possible spatialrequirement and be suitable for the support or creation of bloodpressure, in a manner which as gentle as possible to the blood.Moreover, it should be suitable for covering a large range of the bloodpressure with an as suitable as possible volume flow.

SUMMARY

A blood pump according to the invention comprises a hollow body, inwhich an impeller with a blading is provided, for the production of anaxial propulsion of the blood along the impeller, as well as at leastpartly actively stabilised magnetic bearing device and a hydrodynamicbearing device for the impeller, wherein the impeller may be set into arotation about a rotation axis of the impeller, with a motor statorlocated outside the hollow body, and wherein the hollow body comprisesan inlet for the flow of blood into the hollow body in an inflowdirection essentially parallel to the rotation axis, and an outlet forthe flow of the blood out of the hollow body in a flow-out direction,wherein the outlet is arranged offset with respect to the rotation axisof the impeller for producing an outflow angle between the inflowdirection and the outflow direction, which is different from zero.

Thereby, an outlet is to be understood as an opening in a wall of thehollow body, wherein the outlet as a rule is led further to the outsideby a connection union.

The invention is based on the concept of achieving an as small aspossible spatial requirement of the blood pump, including the necessaryflexible connection tubes, by way of the angle between the inflowdirection of the blood and the outflow direction of the blood beingdifferent from zero, wherein by way of a suitable selection of thisangle, the outlet of the implanted blood pump may be aligned in thedirection of the blood vessel, to which the blood pump is to beconnected, thus for example the aorta, the pulmonary artery stem oranother blood vessel. In this manner, one may select a particularlyshort flexible connection tubing between the blood pump and the bloodvessel, since the flexible connection tubing may be lead in an asstraight-lined as possible manner and in a direct path, to the bloodvessel and not along a curvature forming a detour. Moreover, theapplication of flow bends or elbows for deflecting a flow direction ofthe blood as a rule thus becomes superfluous. This angle preferably liesin a range of between 30° and about 150°, particularly preferably in arange between about 75° and about 105°, wherein with a given angle of90° the blood leaves the blood pump at a right angle to the rotationaxis, and with a given angle of 0° the outflow is effected in the axialdirection.

With the application of conventional axial blood pumps, as a rule,greatly curved flexible connection tubings are necessary, sinceconventional axial blood pumps always have an axial outlet (i.e. theangle between the inflow direction and the outflow direction of theblood is about 0°).

The basic principle of an axial propulsion of the blood by the impelleris retained by the blood pump according to the invention. This isadvantageous, since those blood pumps which are also indicated as axialpumps and which mainly exert an axial force effect on the blood and thusaccelerate this mainly axially, thus in the direction of the rotationaxis along the impeller, deliver the blood in a particularly gentlemanner. In comparison to this, so-called radial pumps accelerate theblood mainly radially to the rotation axis of the impeller. Radial pumpsmoreover mostly have a radial outlet, which often entails the advantageof particularly short connections to blood vessels.

The blood pump according to the invention which is suggested here, thusunifies the advantage of an axial pump with regard to the gentledelivery of blood, as well as the advantage of a radial pump with regardto the shorter flexible connection tubes between the blood pump and theblood vessels and thus with regard to a small spatial requirement aswell as improved rotor dynamics.

For generating the axial propulsion of the blood according to theinvention the blading of the impeller is designed as a spiral (helix).Such a spiral-shaped blading may by single-start or multi-start. Thus,the spiral can comprise one or more, preferably two to six, individualspiral-shaped, i.e. helical, blades (vanes). Each of these blades of theblading may revolve partially, completely or several times around theimpeller with respect to the rotational axis of the impeller.Preferably, the individual blades are wound at least once, morepreferably at least one and a half time around the impeller.Furthermore, the surfaces of the at least one blade of the spiralpointing downstream and the rotational axis of the impeller enclose anon-zero angle (blade angle). The blade angle is related to the pitchand the lead of the blading. Analogously to common screw threadnomenclature, the pitch of the blading is defined as the axial distancebetween two neighboring windings of a blade if the blading is asingle-start blading, and as the axial distance between two neighboringblades if the blading has two or more individual blades. The respectivedistances mentioned above are always measured between surfaces of theblades or blade windings facing in the same axial direction (and notfacing towards each other). Accordingly, the lead of the blading of theimpeller may be defined as the axial distance a volume element of bloodadjacent to a blade of the blading is axially advanced when the impelleris turned one revolution (neglecting any tangential movements of saidvolume element for simplicity). In order to determine the pitch and thelead at the axial ends of the blading, the blading may be extrapolatedin axial direction. In the special case of a constant pitch the lead ofan N-start spiral-shaped blading equals the pitch times N.

By modifying the local blade angle, the pitch and the lead of theblading, the transport effect of the blading on the blood can beadjusted. The blade angle, the pitch as well as the lead may vary alongthe axial extent (length) of the impeller. The blade angle, pitch orlead at a given axial position of the impeller is therefore referred toas local blade angle, local pitch or local lead, respectively.

Preferably, the local pitch of the blading along an entire axial extentof the blading lies in a range between 2 mm und 20 mm, more preferablyin a range between 3 mm und 15 mm. The local pitch of the blading alongthe entire axial extent of the blading preferably lies in a rangebetween 2 mm und 120 mm, more preferably in a range between 3 mm und 40mm. The local blade angle along the entire axial extent of the bladingpreferably lies in a range between 80° and 20°, more preferably in arange between 70° and 30°.

The local pitch and the local lead may increase from an upstream side ofthe blading to a downstream side of the blading. At the upstream side,the local pitch may lie in a range between 2 mm and 8 mm and on thedownstream side between 10 mm and 20 mm. At the upstream side, the locallead may lie in a range between 2 mm and 50 mm and on the downstreamside between 10 mm and 120 mm. Preferably the local pitch and the locallead increase monotonously from the upstream side to the downstream sideof the blading. The blade angle at the upstream side is preferably in arange between 80° and 45°, more preferably between 75° and 55°. Theblade angle at the downstream side of the blading is preferably in arange between 70° and 35°, more preferably between 60° and 40°Preferably the blade angle decreases monotonously from the upstream sideto the downstream side of the blading.

Averaging the local pitch, the local lead and the blade angle along theaxial extent of the blading yields the average pitch, the average leadand the average blade angle, respectively. Preferably, the average pitchlies in a range between 5 mm and 12 mm and the average lead in a rangebetween 5 mm and 85 mm. Moreover, the average blade angle preferablylies in a range between 45° and 65°.

According to the invention, the blood pressure build not only resultsfrom pushing the blood axially forward via the blading, but additionallyalso by transferring a tangential flow velocity and with it rotationalenergy onto the blood. Generally, the tangential flow velocity of theblood and the amount of rotational energy transferred onto the blood viathe blading increases with increasing pitch and lead and with decreasingblade angle.

In the special embodiments shown in the figures below, the bladingextends into the spiral housing of the blood pump. Since the outflow ofblood from the spiral housing (volute) runs tangentially to theimpeller, the above mentioned tangential velocity component of the bloodis efficiently used for building up the pressure.

The individual spiral-shaped blades of the blading are preferablydesigned continuously along the length of the impeller. Furthermore, theblading preferably spreads over at least 80% of the axial extent(length) of the impeller, more preferably over at least 90% of thelength of the impeller and most preferably over the total length of theimpeller. In this way, the impeller is particularly suitable forproducing a gentle, low-turbulence axial propulsion of the blood.

Preferably, an outer contour of the spiral is designed in acylinder-shaped manner. Furthermore, also a peripheral surface of theimpeller, which carries the blading of the impeller between theupstream-side and the downstream-side, may by essentiallycylinder-shaped, truncated cone-shaped or cone-shaped. It can also beenvisaged to vary the height of the blading along the impeller,preferably to increase the height inside the spiral housing. Moreover,the impeller is preferably elongated in direction of the rotationalaxis. Preferably, the impeller has a maximal total diameter (includingthe blading and measured perpendicularly to the axis of rotation) whichis not larger than 60% of the axial extent (length) of the blading ofthe impeller, more preferably not larger than 30% of the axial length ofthe blading of the impeller. An elongated form of the impeller allowsfor a particularly slim shape of the blood pump.

In one embodiment, it is envisaged that a maximal radial extent of theblading, i.e. a maximal height of the blading, is smaller than 50% of amaximal total radius of the impeller (measured perpendicularly to theaxis of rotation and including the blading), preferably the maximalheight of the blading is smaller than 30% of the maximal total radius ofthe impeller. Typically, the maximal height of the blading lies in arange between 1 mm und 4 mm, more preferably in a range between 1.5 mmund 3 mm.

The at least one blade of the blading may have a maximal width (measuredperpendicularly to the rotational axis and perpendicularly to the heightof the blading) which is smaller than 10% of a maximal totalcircumference of the impeller (measured perpendicularly to the axis ofrotation and including the blading), preferably the maximal width isless than 5% of the maximal total circumference of the impeller.Typically, the maximal width lies in a range between 0.5 mm and 3 mm,preferably between 1 mm und 2 mm. In this manner, the at least one bladehas the form of thin helical rib.

The blood pump according to the invention is further characterised bythe omission of an outlet guide vane which is mounted downstream of theimpeller in the flow direction. Such a downstream guide vane in commonaxial pumps with an axial outlet, serves for converting rotationmovements of the blood into an additional axial pressure build-up andthus for an efficiency increase of the axial delivery of the blood. Withthe non-axial outlet according to the invention, the rotational movementof the blood at least partly also contributes to the blood pressurewhich is produced by the blood pump, which may advantageously also beutilised by way of omitting the downstream guide vane. Moreover, themechanical loading of the blood by way of a deflection of the blood bythe downstream guide vane is avoided by way of omission of thedownstream guide vane, by which means the danger of damage to the bloodis further reduced.

A further significant advantage which is achieved by way of omitting thedownstream guide vane lies in a smaller axial length of the blood pumpand thus in a reduced spatial requirement of the blood pump.

Moreover, the problem of unfavourable onflow angles of the downstreamguide vane at certain operating points of the blood pump is also doneaway with due to the omission of the downstream guide vane. Ifspecifically the downstream guide vane is subjected to onflow at anunfavourable angle, then even pressure losses may arise due to thedownstream guide vane. Moreover, local pressure fluctuations may beformed at the downstream guide vane, which have an unfavourable effecton the flow course of the blood on the impeller and render a stablebearing of the impeller more difficult, in particular with lowerrotational speeds.

An at least partly actively stabilised magnetic bearing device for theimpeller, as is described for example in WO 00/64030, is basicallysuitable for the contact-free accommodation of radial as well as axialforces. In particular, an active stabilisation of the axial bearing(active axial stabilisation) has been found to be particularlyadvantageous with all rotational speeds. The magnetic bearing device maycomprise permanent-magnetic elements integrated into the impeller forthe purpose of an active axial stabilisation (the impeller as a rulealso contains permanent-magnetic elements for the functioning as a motorrotor of an electric motor). Additionally, the magnetic bearing systemmay comprise ring coils for an active axial stabilisation, which permitan active stabilisation (closed-loop control) of the axial position ofthe impeller by way of an axial magnet flux. These ring coils areindependent of a motor winding and serve exclusively for the activelystabilised, axial bearing of the impeller. The mentioned ring coils mayfor example be arranged outside the hollow body such that they surroundthese in an annular manner. Moreover, the magnetic bearing device maycomprise a sensor system for measuring a position of the impeller, inparticular for ascertaining a deviation from an axial desired position,as well as a closed-loop control unit which is connected to the sensorsystem and the ring magnets and which sets the magnetic flux produced bythe ring magnets, according to the measured axial position of theimpeller, for correcting a possible deviation of the impeller from thedesired position. Further details are to be deduced for example from theabove mentioned document or the description of special embodiments ofthe invention further below.

In a further development, the magnetic bearing device of the blood pumpsmoreover comprises permanent-magnetic bearing elements for a passiveradial bearing (passive radial stabilisation) of the impeller. Thesepermanent-magnetic bearing elements may for example be arranged withinthe hollow body in the direct vicinity to an upstream-side or adownstream-side of the impeller, for example in a hub or the impeller, aguide vane or a termination plate of the hollow body. Further detailsare to be deduced again from the mentioned document or the descriptionof special embodiments of the invention further below.

The hydrodynamic bearing of impellers in blood pumps is basically known.In one embodiment of the invention, one envisages the hydrodynamicbearing device of the impeller as in WO 02/66837 being realised by asupport ring connected to the impeller or several such support ringsand, in this manner, forming an annular gap (or several annular gaps)between the support ring and an inner wall of the hollow body, for aradial bearing of the impeller. Such a support ring, preferably formedas a rotationally-symmetrical hollow cylinder, may be designed indifferent widths and may be fastened on the impeller at any location, inorder to achieve an optimal stabilisation of the impeller, in particularwith respect to tilting of the impeller. In this manner, one maycompensate hydrodynamic and mechanical imbalances of the impeller in aparticularly effective manner. This is particularly advantageous, if asis described further below, a (spiral-shaped) discharge channel ispartly peripheral around the impeller. In this case, a suitable supportring directly upstream of the discharge channel in the flow direction,may contribute to a stabilisation of the impeller.

A further development of the invention envisages the outlet of thehollow body being arranged between an upstream-side of the impeller,said upstream-side being away from the inlet, and a downstream-side ofthe impeller, said downstream-side being away from the inlet. In thismanner, one may realise particularly compact embodiments with a reducedconstruction length. Moreover, it has been found that this contributesto particularly good flow characteristics of the blood pump and thusalso to a stabilisation of the impeller, by which means the pressureregion which may be covered by the blood pump is increased. Preferably,the outlet is arranged in a direct environment of the downstream-side ofthe impeller, in order to utilise as much as possible the actualpropulsion due to the blading of the impeller which is preferably bladedover an entire length of its peripheral surface. In order to increasethis, the blading of the impeller preferably also extends along theoutlet, i.e. at the height of the outlet.

In a further development, the blood pump comprises a backing plate whichis arranged essentially perpendicular to the rotation axis of theimpeller, for terminating the hollow body. In the case that the outlet,as described above, is arranged between the upstream-side and thedownstream-side of the impeller, this backing plate is preferablyarranged in the direct vicinity of the downstream-side of the impeller,in order thus to avoid flow-free dead spaces between the impeller andthe backing plate. Such dead spaces generally entail an increased riskof blood clots and are therefore to be avoided as much as possible. In afurther development, the backing plate is designed as a closure which iseasy to open, for the simple creation of an axial access into the hollowbody, such as for a simpler assembly or an adjustment of the blood pumpduring the assembly process.

In one embodiment, one envisages an inner radius of the hollow bodybeing enlarged for forming a discharge channel which runs tangentiallyaround the impeller and runs into the outlet (spiral housing). Such adischarge channel permits a discharge of the blood through the outletout of the hollow body, essentially tangential to the impeller (moreprecisely tangentially to a peripheral surface of the impeller). In thismanner, a respective tangential flow component of the blood and thekinetic energy of the blood which this entails, are retainedparticularly well and in a low-loss manner on flowing out of the hollowbody and may be utilised for the efficient production of the bloodpressure. A flow component of the blood which runs tangentially to theperipheral surface but perpendicularly to the rotation axis, inprinciple always arises with axial pumps, since the propulsion producedby the impeller, apart from the axial component, also has a componentperpendicular to the rotation axis. In particular, eddies withcorresponding energy losses as may occur with a simple radial outlet,are avoided by way of such a discharge channel. Moreover, in thismanner, the flow characteristics of the blood pump are improved and thusalso the stability of the impeller, by which means the pressure rangewhich may be covered by the blood pump is further increased. Moreoversimultaneously, the mechanical loading of the blood which the mentionededdies entail, may be largely avoided.

In order to achieve an as high as possible efficiency of the blooddelivery with an as compact as possible construction shape, the bladingof the impeller also extends along the discharge channel, i.e. at aheight of the discharge channel.

A further development envisages the discharge channel being widenedtowards the outlet and thus being designed in a spiral manner like aspiral housing. In this manner, a continuous reduction of the flow speed(with a constant volume flow) towards the outlet, and a reduction ofeddy formations is achieved, and thus a particularly gentle flow of theblood out of the hollow body. Moreover, the pressure increase producedby the blood pump, at the outlet in the transition to the outletchannel, may be transmitted in a particularly effective manner by way ofthe reduced flow speed of the blood. Typical flow speeds at the outletof the blood pump in this example lie of the magnitude below 1 m/s.

In a further embodiment, one envisages a peripheral surface of theimpeller, which carries the blading of the impeller between theupstream-side and the downstream-side, being essentiallycylinder-shaped, truncated cone-shaped or cone-shaped, for theproduction of as uniform and eddy-free as possible propulsion. In thismanner, a mainly axial propulsion is ensured, thus a propulsion which isgentle with regard to the blood. However, one may also achieve anadditional radial acceleration component by way of a diameter of theimpeller which increases towards the downstream-side of the impeller.

In particular, in combination with a cylinder-shaped ortruncated-cone-shaped design of the impeller, one may envisage an inletguide vane which is arranged on the upstream-side of the impeller. Thisinlet guide vane on the one hand serves for an as eddy-free as possibleonflow of the impeller, thus an onflow which is gentle to the blood andis as loss-free as possible, and may furthermore comprise a stationaryblading, in order to reduce the rotation movement of the blood about therotation axis of the impeller and convert it into an axial propulsionfor further increasing the delivery output of the blood pump.Preferably, the inlet guide vane is arranged in the direct vicinity tothe upstream-side of the impeller, for avoiding or reducing a non-flowdead space between the inlet guide vane and the impeller. Moreover, theinlet guide vane, as already described, may contain permanent-magneticbearing elements which are components of the magnetic bearing device.

In a further development, one envisages the blood pump comprising acontrol unit which is set up to set rotational speeds of the impeller ina range between 3000 rpm and 35000 rpm, for producing a blood pressureat the outlet in a range between 5 mmHg and 150 mmHg with a volume flowof the blood which is adapted to requirements. In this manner, a volumeflow between 0 l/min and 20 l/min may be set depending on the flowresistance of the pulmonary circulation or the systemic circulation.

An alternative blood pump comprises a hollow body, in which an impelleris provided with a blading, for producing an axial propulsion of theblood along the impeller, wherein the impeller may be set into rotationabout a rotation axis of the impeller with a motor stator which islocated outside the hollow body, and wherein the hollow body comprisesan inlet for the flow of blood into the hollow body in an inflowdirection which is essentially parallel to the rotation axis, and anoutlet for the flow of the blood out of the hollow body in an outflowdirection, wherein the outlet is arranged offset with respect to therotation axis of the impeller, for producing an outflow angle betweenthe inflow direction and the outflow direction, which is different fromzero, wherein an inner radius of the hollow body is enlarged for forminga spiral-shaped discharge channel which runs tangentially around theimpeller and which runs out into the outlet, for the flow of the bloodout of the hollow body, said flow running essentially tangentially tothe impeller, wherein the outlet of the hollow body is arranged betweenan upstream-side of the impeller, said upstream-side facing the inlet,and a downstream-side of the impeller, said downstream-side being awayfrom the inlet.

Such a blood pump may contain a mechanical, hydrodynamic, a magnetic ora hybrid bearing device, for bearing the impeller. Moreover preferably aperipheral surface of the impeller, which carries the blading, may bedesigned in an essentially cylinder-shaped manner for an axialpropulsion of the blood.

All technical features which are specified and described above are to beconsidered for the further development of this alternative blood pump.The mentioned advantages arise in each case. For the sake ofcompleteness, the features are one again cited in a brief manner. Theembodiments further above are referred to for a more detailedexplanation.

Thus the hydrodynamic bearing device of the impeller may be designed asa support ring connected to the impeller, for the formation of anannular gap between the support ring and an inner wall of the hollowbody. It is also possible for the outlet of the hollow body to bearranged between an upstream-side of the impeller, said upstream-sidefacing the inlet, and a downstream-side of the impeller, saiddownstream-side being away from the inlet. One may also envisage aninner radius of the hollow body being enlarged for the formation of adischarge channel which runs tangentially around the impeller and runsout into the outlet, for the flowing-away of the blood out of the hollowbody, which is essentially tangential to the impeller. Moreover, it ispossible for this discharge channel to widen towards the outlet.

Moreover, one may envisage the magnetic bearing device comprising anactively stabilised axial bearing.

Furthermore, it is possible for a peripheral surface of the impeller,which carries the blading, being designed in a cylinder-shaped,cone-shaped or truncated-cone-shaped manner. The blading of the impellermay be designed as a spiral. Also all features concerning the shape ofthe blading, the blades of the blading and of the impeller as describedabove may be realized with this alternative blood pump.

Moreover, one may envisage an inlet guide vane, which may include partsof the magnetic bearing device.

Finally, the blood pump may comprise a control unit which is set up toset rotational speeds of the impeller in a range of between 3000 rpm and35000 rpm for producing a blood pressure at the outlet in a rangebetween 5 mmHg and 150 mmHg with a volume flow which is adapted tophysiological conditions.

In a total artificial heart according to the invention, one envisagesproviding two blood pumps of a type suggested here, wherein a firstblood pump is preferably used as an LVAD and a second blood pump as anRVAD. The total artificial heart is particularly space-saving and maythus be arranged in the thorax at the heart in a particularly simplemanner by way of the application of the blood pumps described here.

In one embodiment, one envisages the impellers of both blood pumps ofthe total artificial heart being arranged on a common rotation axis, bywhich means a particularly simple design and assembly is made possible.Moreover, this permits an advantageous slim shape of the totalartificial heart, by which means an implantation in the thorax issimplified.

In one embodiment example, the impellers of both blood pumps are fixedlyconnected to in another into a single, common impeller, wherein thecavities of both blood pumps are grouped together into a common hollowbody (housing). This permits a construction of the total artificialheart pump which is particularly short axially. Moreover, a simplebearing is possible in this manner, since the common impeller has lessdegrees of freedom than two individual impellers.

In a further embodiment example, one envisages a bearing block beingpresent between the impeller of the first blood pump and the impeller ofthe second pump, in which bearing block at least parts of the bearingdevice of the first and/or second blood pump are integrated.

BRIEF DESCRIPTION OF THE DRAWINGS

Special embodiments of the invention are described hereinafter in moredetail by way of FIGS. 1A, 1B, 2-6. The same reference numerals indicatethe same features with regard to subject-matter. There are shown in:

FIG. 1A a schematic representation of a longitudinal section through ablood pump of the type suggested here,

FIG. 1B a schematic representation of a longitudinal section through ablood pump of the type suggested here,

FIG. 2 a schematic representation of a longitudinal section through ablood pump of the type suggested here,

FIG. 3 a schematic representation of a cross section through a hollowbody of a blood pump of the type suggested here,

FIG. 4 a schematic representation of a partly cutaway hollow body of ablood pump of the type suggested here,

FIG. 5 a schematic representation of a total artificial heart of thetype suggested here, with a single impeller and

FIG. 6 a schematic representation of a total heart pump of the typesuggested here, with two individual impellers.

DESCRIPTION OF PREFERRED EMBODIMENTS

A schematic representation of a longitudinal section through a bloodpump 1 of the type suggested here is represented schematically in FIG.1A. The blood pump 1 comprises a hollow body 2 (represented as acontinuous thick line), in which an impeller 3 with a blading 4 isprovided. Moreover, the hollow body 2 comprises an inlet 5 for the flowof blood in an inflow direction which is parallel to a rotation axis R(shown dashed), and an outlet 6 for the outflow of blood in an outflowdirection which runs perpendicular to the section plane. Accordingly, inthis embodiment example, the outlet is arranged offset at a right anglerelative to the rotation axis R, for producing an outflow angle α ofα=90°, which is different from zero, between the inflow direction andthe outflow direction.

The outlet 6 of the hollow body 2 is arranged between an upstream-side 9of the impeller 3, said upstream-side facing the inlet, and adownstream-side 10 of the impeller 3, said downstream-side being awayfrom the inlet. An inner radius of the hollow body 2 serves for forminga discharge channel 11 which runs tangentially around to the impeller 3and runs out into the outlet 6, for a discharge of the blood out of thehollow body 2, said discharge running essentially tangentially to theimpeller 3.

Moreover, a hydrodynamic bearing device is provided which is designed astwo support rings 7 connected to the impeller 3, for the formation oftwo annular gaps 8 between the support rings 7 and an inner wall of thehollow body 2, for a radial bearing of the impeller 3.

A peripheral surface 12 of the impeller 3, which carries the blading 4,is formed in a cylinder-shaped manner, but may just as well be designedin a truncated-cone-shaped or cone-shaped manner. The axial dimension(length) L of the impeller is selected larger than a diameter D of theimpeller on the downstream-side of the impeller. The blading of theimpeller is characterised by a pitch which increases towards the outlet6. In this manner one permits an axial propulsion up to the dischargechannel 11, which is particularly gentle to the blood. The blading ofthe impeller 4 extends axially completely (in other embodiments partlyor not at all) into the discharge channel 11 and the outflow 6.

An inlet guide vane 14 which is provided with a blading 14′, is providedin the direct vicinity of the upstream-side 9 of the impeller 3.

The blood pump further comprises a partly actively stabilised bearingdevice which contains an actively stabilised, magnetic axial bearing aswell as a passive, magnetic radial bearing. The magnetic bearing devicefirstly comprises two permanent magnets 15, 15′ which are arranged inthe impeller at the upstream-side and at the downstream-side.Furthermore, two further permanent magnet bearings 16, 16′ which arepoled opposite to these (attracting) and which are integrated into theinlet guide vane 14 and the backing plate 13, respectively, serve theformation of the passive, magnetic radial bearing, which ensures thatthe impeller 3 is held in a radial desired position between the inletguide vane 14 and the backing plate 13. Moreover, for the activelystabilised magnetic axial bearing, two ring coils 17, 17′ are arrangedoutside the hollow body 2, in front of and behind the impeller 3, suchthat they are peripheral around the hollow body 2 in an annular mannerfor producing an axial magnetic flux. Moreover, the magnetic bearingdevice comprises a sensor system which comprises distance sensors 18,18′ integrated into the inlet guide vane 14 and/or the backing plate 13as well as into the impeller 3, for measuring the gap widths between theimpeller 3 and the inlet guide vane 14 or the backing plate 13, as wellas a closed-loop control unit (not shown here) which is connected to thedistance sensors 18, 18′ and the ring magnets, said closed-loop controlunit setting the magnet flux produced by the ring magnets, according tothe measured axial position of the impeller, for correcting a possibledeviation of the impeller from an axial desired position.

Finally, a motor winding 19 running around the hollow body and a motormagnet 20 integrated into the impeller are provided, said motor magnetbeing magnetised in an alternating radial manner, for driving theimpeller 2.

In FIG. 1B, a schematic representation of a longitudinal section througha blood pump 1 of the type suggested here is shown, which differs fromthe blood pump described by way of FIG. 1A in that a central,cylindrical rod 16 extends from a downstream-side 27 of the pump 1axially into the hollow body 2 towards the impeller 3. In said rod 26,one of the distance sensors 18′ is integrated for measuring the gapwidth between the impeller 3 and the rod 26 as well as one of thepermanent magnet bearings 16′ being a part of the passive, magneticradial bearing. Furthermore, the ring coil 17′ of the activelystabilized axial bearing now is positioned axially before the outlet 6and runs around the hollow body 2, while in the embodiment shown in FIG.1A, the respective ring coil 17′ is located behind the hollow body 2(with respect to the axial pump direction) and consequently does not runaround the hollow body 2. All other features of the embodiment shown inFIG. 1B are identical to the features of the embodiment shown in FIG.1A.

In FIG. 2, a schematic representation of a longitudinal section througha blood pump 1 of the type suggested here is shown, which differs fromthe blood pump described by way of FIG. 1A only by way of a changedhydrodynamic bearing device. In the example shown in FIG. 2, this isdesigned as a single support ring 7 connected to the impeller 3, forforming an individual annular gap 8 between the support ring 7 and aninner wall of the hollow body 2, for a radial bearing of the impeller 3

Moreover, a radius r of the hollow body 2 at a height of the dischargechannel 11 is represented, wherein this radius increases towards theoutlet, for forming a spiral-shaped discharge channel 11 which widenstowards the outlet. A radius of the impeller blading is indicated as r′.It is the case that r′<r.

For the embodiment shown, it is r′=8 mm and r=14 mm. Furthermore, theimpeller is elongated, having an axial extent (length) of 40 mm. Theblading is spread over the entire length l of the impeller 3 so that theaxial extent of the blading 4 is also 40 mm. The maximal total diameterof the impeller 3 is given by 2r′=16 mm, which is less than 50% of theaxial blading extent.

The blading 4 of the impeller 3 consists of 2 blades 4 (two-startblading) each of them having a maximal height of 2 mm, which is lessthan 30% of the maximal total impeller radius r′. The maximal width ofthe blades 4 is 1.5 mm which is less than 5% of the maximal totalcircumference of the impeller 3 (given by 2□□r′=52.27 mm). Moreover, theblades 4 each run about 1.8 times around the impeller 3 (with respect tothe rotational axis R).

At the upstream-side 9 of the impeller 3, a local pitch of the blading 4is about 5 mm and a local lead is about 12 mm. The local pitch and thelocal lead monotonously increase towards the downstream-side 10 of theimpeller 3 to a pitch value of about 12 mm and a lead value of about 40mm at the downstream-side 10 of the impeller 3, respectively. On averagethe pitch is about 10 mm and the lead is about 30 mm. At theupstream-side 9 of the impeller 3, a blade angle of the blades is about75° and monotonously increases towards the end 10 of the impeller to avalue of about 45°. On average the blade angle is about 60°.

Note that the explicit values given above for quantifying the design ofthe impeller, the blading and other parts of the blood pump shown in thefigures only serve for illustrative purposes and by no means arerestrictive. All parts of the blood pump can be modelled and reshaped toachieve desired pump characteristics. Preferred ranges for variousparameters of the pump design are given further above in the generalpart of the description.

FIG. 3 shows a schematic representation of a cross section through ahollow body 2 of a blood pump 1 according to FIG. 1A or FIG. 2. Thecross section runs perpendicularly to the rotation axis R through thedischarge channel 11 of the hollow body 2 of the blood pump 1. Thehollow body 2 has a radius r which in comparison to a radius r′ of thehollow body is increased to a height of the upstream-side 3 of theimpeller 3, for forming the discharge channel 11. The discharge channel11 widens in a spiral-like manner in its course towards the discharge 6and in this manner forms a spiral housing. The discharge 6 is continuedto the outside into a connection union 21, which is widened further tothe outside for reducing the flow speed of the blood.

FIG. 4 shows a schematic representation of a partly cutaway hollow body2 of a blood pump 1 according to FIG. 3. Again, one may recognise thehollow body 2 with the inlet 5 for the inflow of blood in an inflowdirection which is indicated by the arrow indicated at E, with atangential outlet 6 which is lengthened into an outlet union 21, for theoutflow of the blood in an outflow direction which is indicated by thearrow indicated at A and which runs at a right angle to the inflowdirection E.

The cylinder-shaped axial impeller 3 is arranged in the hollow body,wherein FIG. 4 additionally, by way of example, shows the covering of apart of the impeller by the spiral-shaped outlet channel. Thespiral-shaped outlet channel 11 runs tangentially to the impeller 3,runs out into the outlet 6 and in this manner forms a spiral chamber(spiral housing).

One embodiment of a total heart pump 22 of the type suggested here isschematically represented in FIG. 5. It comprises two blood pumps 1, 1′of the type suggested here, whose hollow bodies 2, 2′ are connectedaxially into a common hollow body. This at its two ends comprises twoinlets 5, 5′ for the inflow of blood from the pulmonary circulation orthe systemic circulation, so that the right blood pump 1 is envisaged asan RVAD and the left blood pump 1′ as an LVAD. The two impellers 3, 3′of the two blood pumps 1, 1′ are axially connected to one another in afixed manner into a common impeller. The blood may be driven in an axialmanner by way of a suitable design of the blading 4, 4′ of the commonimpeller 3, 3′, towards a middle of the common hollow body 2, 2′, atwhich two spiral-shaped outlet channels 11, 11′ (spiral chambers) areformed, which in each case run out into an outlet 6, 6′ for thetangential (right-angled) outflow of the blood out of the common hollowbody 2, 2′.

The blading 4, 4′ of the impeller is designed for producing twodifferent values of the blood pressure at the two outlets 6, 6′. Thepitch of the spiral-shaped blading is correspondingly adapted for thispurpose.

The design parameters of the left pump 1′, in particular defining theshape of the impeller 3′ and the blading 4′ are equal to the designparameters of the blood pumps shown in FIGS. 1-4. The right blood pump1, however, has an opposite handedness and, moreover, smaller pitch andlead values in order to produce smaller blood pressure values than theleft blood pump 1′ at same rotational frequency. All other parametersare the same as for the left blood pump 1′. In this example, at theupstream-side of the impeller 3 a local pitch of the blading is about 3mm and a local lead is about 10 mm. The local pitch and the local leadmonotonously increase towards the downstream-side of the impeller 3 to apitch value of about 8 mm and a lead value of about 25 mm at thedownstream-side of the impeller 3, respectively. On average the pitch isabout 5 mm and the lead is about 17 mm. At the upstream-side of theimpeller 3, a blade angle of the blades is about 80° and monotonouslyincreases towards the end 10 of the impeller to a value of about 55°. Onaverage the blade angle is about 65°.

A connection gap 23 between the common hollow body 2, 2′ and the commonimpeller 3, 3′ exists between the two outlet channels 11, 11′. Theconnection gap 23 may be designed as narrowly as possible in order toreduce a leakage flow of the blood between the cavities 3, 3′ of thefirst and the second blood pump 1, 1′.

Moreover, the total artificial heart, at the two inlets, in each casecomprises an olive 24, 24′ (connection piece) for the connection of aflexible connection tubing.

One embodiment of a total artificial heart 2 of the type suggested here,is schematically represented in FIG. 6. It comprises two blood pumps 1,1′ of the type suggested here, whose hollow bodies (cavities) 2, 2′ arealigned axially to one another and are connected to one another in afixed manner via a bearing block 25. The bearing block contains parts ofthe bearing devices (e.g. permanent-magnetic bearing magnets for theaxial bearing) of the two blood pumps 1, 1′ for bearing the twoimpellers 3, 3′. These are not mechanically connected to one another andthus may be rotated about the common rotation axis R independently ofone another. The two inlets 5, 5′ are envisaged for the flow of bloodfrom the pulmonary circulation or the systemic circulation, so that asin the previous embodiment example, the right blood pump 1 is envisagedas an RVAD and the left blood pump 1′ as an LVAD. The blood may beaxially driven in the direction of the bearing block 25 by way of asuitable choice of the rotational speed and/or by way of a differentdesign of the blading 4, 4′ of the two impellers 3, 3′.

The design parameters of the left pump 1′ and the right pump 1, inparticular defining the shape of the impeller 3′ and 3 and the blading4′ and 4 are equal to the design parameters of the blood pumps shown inFIGS. 1-4.

Moreover, spiral-shaped outlet channels 11, 11′ (spiral chambers) areprovided in each case at the downstream-sides 10, 10′ of the twoimpellers 3, 3′ and these outlet channels in each case run out into anoutlet 6, 6′ for the tangential (right-angled) flow of blood out of thecavities 2, 2′.

As described above and shown in FIGS. 5 and 6, the downstream-sides oftthe impellers 3, 3′ of the total artificial hearts 22 shown are facingtowards each other so that the blood is pumped towards a center of thetotal artificial hearts 22 located between the impellers 3 and 3′, i.e.the blood is pumped towards the connection gap 23 (FIG. 5) or towardsthe bearing block 25 located between the two hollow bodies 2, 2′ (FIG.6). So the orientations of the axial propulsion of the two blood pumps1, 1′ of the total artificial hearts 22 are anti-parallel and directedtowards each other.

The invention claimed is:
 1. A blood pump, comprising: a hollow body; amotor stator located outside the hollow body; an inlet stator; an outletstator; and an impeller including a blading, wherein the impeller ispositioned in the hollow body, the impeller is configured to produce anaxial propulsion of blood along the impeller in an axial direction, andthe blading has a pitch that increases along the impeller along theaxial direction, wherein the motor stator is configured to cause theimpeller to rotate about a rotation axis of the impeller, wherein thehollow body comprises an inlet for a flow of blood into the hollow bodyin an inflow direction that is essentially parallel to the rotationaxis, and an outlet for an outflow of blood out of the hollow body in anoutflow direction, and wherein the outlet is arranged offset to therotation axis of the impeller, the outlet is configured to produce anoutflow angle between the inflow direction and the outflow direction,and the outflow angle is non-zero; and wherein the impeller is arrangedbetween the inlet stator and the outlet stator in the axial direction,and the outlet stator not being held in the flow path by diffuserblades.
 2. A total artificial heart comprising two blood pumps accordingto claim
 1. 3. The total artificial heart of claim 2, wherein theimpellers of the two blood pumps are arranged on a common rotation axis.4. The total artificial heart of claim 2, wherein orientations of theaxial propulsion of the two blood pumps are anti-parallel and directedtowards each other.
 5. The blood pump of claim 1, wherein the bladingdoes not extend up to an outlet side end of the impeller.
 6. The bloodpump of claim 1, wherein the outlet of the hollow body is arrangedbetween an upstream-side of the impeller, said upstream-side facing theinlet, and a downstream-side of the impeller, said downstream-side beingaway from the inlet.
 7. The blood pump of claim 1, wherein an innerradius of the hollow body is enlarged for forming a discharge channelwhich runs tangentially around the impeller and runs out into theoutlet, for a flowing-away of the blood out of the hollow body, runningessentially tangentially to the impeller.
 8. The blood pump of claim 1,wherein the discharge channel widens towards the outlet.
 9. The bloodpump of claim 1, wherein the blading comprises at least two blades andthe pitch is measured along the axial direction between adjacent blades.10. The blood pump of claim 1, wherein a peripheral surface of theimpeller, said peripheral surface carrying the blading, is designed inan essentially cylinder-shaped manner, cone-shaped manner ortruncated-cone-shaped manner.
 11. The blood pump of claim 1, wherein thepitch of the blading lies in a range between 2 mm and 20 mm along theaxial direction of the blading.
 12. The blood pump of claim 1, whereinthe pitch of the blading at an inlet side beginning of the blading liesin a range between 2 mm and 8 mm and the pitch of the blading at anoutlet-side end of the blading lies in a range between 10 mm and 20 mm.13. The blood pump of claim 1, wherein the blading comprises at leastone spiral-shaped blade which is wound at least once around theimpeller.
 14. The blood pump of claim 1, wherein a maximal height of theblading is less than 50% of a maximal total radius of the impeller. 15.The blood pump of claim 1, wherein a maximal width of the blading isless than 10% of a maximal total circumference of the impeller.
 16. Theblood pump of claim 1, wherein the blading is spread over at least 80%of an axial length of the impeller.
 17. The blood pump of claim 1,wherein the impeller has a maximal total diameter which is not largerthan 60% of the total axial extent of the blading of the impeller. 18.The blood pump of claim 1, wherein an inlet guide vane is provided.